Mode-locked femtosecond fiber lasers at 1030 nm and 1550 nm have been improving significantly in the last several years, particularly with respect to the achievable output pulse energy (increasing from 1 to ˜10 nJ). Even higher pulse energy can be achieved in femtosecond fiber sources based on fiber chirped pulse amplification. However, femtosecond fiber sources, including lasers, have seen only limited applications in multiphoton imaging. The main reason is that they offer very limited wavelength tunability (tens of nanometer at best), severely restricting the applicability of these lasers, making them only suitable for some special purposes. In addition, existing femtosecond fiber sources at high pulse energy (>1 nJ) are not truly “all fiber,” i.e., the output is not delivered through a single mode optical fiber. Thus, additional setup (typically involving free-space optics) must be used to deliver the pulses to imaging apparatus, partially negating the advantages of the fiber source.
Reports have demonstrated the possibility of propagating femtosecond IR pulses through a large core optical fiber at intensities high enough (˜1 nJ) for multiphoton imaging. In addition, a special HOM fiber that is capable of delivering energetic femtosecond pulses (˜1 nJ) has been demonstrated. However, both of these fibers have normal dispersion, and both require a free-space grating pair for dispersion compensation. Not only is such a grating pair lossy and complicated to align, it needs careful adjustment for varying fiber length, output wavelength, and output pulse energy, and falls short of the requirement for most biomedical research labs and future clinical applications.
Femtosecond fiber sources are highly robust and cost effective. However, energetic femtosecond fiber sources at 1300 nm prove to be very difficult to obtain because of the characteristics of the fluoride fiber based gain medium. Such difficulties are exemplified by the huge performance and cost gap between optical amplifiers (essentially lasers without the cavity mirrors) at these wavelengths. For example, fiber amplifiers at 1300 nm are expensive (˜$31 k) and have limited output power (˜60 mW) and narrow spectral bandwidth (˜20 nm). This is in sharp contrast to fiber amplifiers at 1030 nm and 1550 nm, where high power amplifiers (˜2 W) will only cost ˜$25 k and has a spectral bandwidth of ˜40 nm. Thus, the route that follows the development of fiber sources at 1030 and 1550 nm is unlikely to be productive at 1300 nm. An alternative approach must be taken in order to create a femtosecond fiber source at 1300 nm that has a comparable performance and cost to that of the fiber sources at 1030 and 1550 nm.
Higher-order-mode (HOM) fiber has attracted significant interest recently, due to the freedom it provides to design unique dispersion characteristics in all-solid (i.e., non-“holey”) silica fiber.
The ‘wavelength tunability’ of femtosecond optical sources has been extensively studied within the phenomenon of soliton self-frequency shift (SSFS), in which Raman self-pumping continuously transfers energy from higher to lower frequencies within an optical fiber. SSFS has been exploited over the last decade in order to fabricate widely frequency-tunable, femtosecond pulse sources with fiber delivery. Since anomalous (positive) dispersion (β2<0 or D>0) is required for the generation and maintenance of solitons, early sources that made use of SSFS for wavelength tuning were restricted to wavelength regimes>1300 nm, where conventional silica fibers naturally exhibit positive dispersion.
In addition, Cherenkov radiation has been demonstrated in microstructured fibers pumped near their zero-dispersion wavelength. In general, an ideal soliton requires a perfect balance between dispersion and nonlinearity so that energy becomes endlessly confined to a discrete packet—both spectrally and temporally. When perturbations are introduced, this stable solution breaks down, allowing the transfer of energy between the soliton and the disturbance. Such energy transfer occurs most efficiently in fibers for solitons near the zero-dispersion wavelength. The spectral regime to which energy couples most efficiently has been dubbed “Cherenkov radiation” due to an analogous phase matching condition in particle physics. The phenomenon of Cherenkov radiation in fibers is often associated with SSFS as it allows a convenient mechanism for more efficient energy transfer between the soliton and the Cherenkov band. In particular, when the third-order dispersion is negative, SSFS will shift the center frequency of the soliton toward the zero-dispersion wavelength, resulting in efficient energy transfer into the Cherenkov radiation in the normal dispersion regime. The problem of tunability remains an issue for these arrangements capable of creating Cherenkov radiation.
The recent development of index-guided photonic crystal fibers (PCF) and air-core photonic band-gap fibers (PBGF) have relaxed this tunability requirement somewhat, with the ability to design large positive waveguide dispersion and therefore large positive net dispersion in optical fibers at nearly any desired wavelength. This development has allowed for a number of demonstrations of tunable SSFS sources supporting input wavelengths as low as 800 nm in the anomalous dispersion regime.
Unfortunately, the pulse energy required to support stable Raman-shifted solitons below 1300 nm in index-guided PCFs and air-core PBGFs is either on the very low side, a fraction of a nJ for silica-core PCFs, or on the very high side, greater than 100 nJ (requiring an input from an amplified optical system) for air-core PBFGs. The low-energy limit is due to high nonlinearity in the PCF. In order to generate large positive waveguide dispersion to overcome the negative dispersion of the material, the effective area of the fiber core must be reduced. For positive total dispersion at wavelengths less than 1300 nm, this corresponds to an effective area, Aeff, of 2-5 μm2, approximately an order of magnitude less than conventional single mode fiber (SMF). The high-energy limit is due to low nonlinearity in the air-core PBGF where the nonlinear index, n2, of air is roughly 1000 times less than that of silica. These extreme ends of nonlinearity dictate the required pulse energy (U) for soliton propagation, which scales as U□D·Aeff/n2. In fact, most microstructure fibers and tapered fibers with positive dispersion are intentionally designed to demonstrate nonlinear optical effects at the lowest possible pulse energy, while air-core PBGFs are often used for applications that require linear propagation, such as pulse delivery.
There are a number of biomedical applications that require femtosecond sources. Although applications requiring a large spectral bandwidth (such as optical coherency tomography) can also be performed using incoherent sources such as superluminscent diodes, techniques based on nonlinear optical effects, such as multiphoton microscopy and endoscopy, almost universally require the high peak power generated by a femtosecond source.
Molecular two-photon excitation (2PE) was theoretically predicted by Maria Goppert-Mayer in 1931. The first experimental demonstration of two-photon absorption, however, came nearly 30 years later, after the technological breakthrough of the invention of the ruby laser in 1960. It was almost another 30 years before the practical application of 2PE for biological imaging was demonstrated at Cornell University in 1990. Once again, this new development was propelled in large part by the rapid technological advances in mode-locked femtosecond lasers. Since then, two-photon laser scanning microscopy has been increasingly applied to cell biology and neurosciences. A number of variations, including three-photon excitation (3PE), second and third harmonic generation imaging, near-field enhanced multiphoton excitation and multiphoton endoscopic imaging, have emerged and further broadened the field, which is currently known as multiphoton microscopy (MPM). Today, MPM is an indispensable tool in biological imaging. Like any nonlinear process, however, multiphoton excitation requires high peak intensities, typically 0.1 to 1 TW/cm2 (TW=1012 W). Besides tight spatial focusing, MPM typically requires pulsed excitation sources to provide additional temporal “focusing” so that efficient multiphoton excitation can be obtained at low average power. For example, a femtosecond laser with 100-fs pulse width (τ) at 100 MHz pulse repetition rate (f) will enhance the excitation probability of 2PE by a factor of 105, i.e., the inverse of the duty cycle (fτ). The development of multiphoton imaging depends critically on ultrafast technologies, particularly pulsed excitation source.
Endoscopes play an important role in medical diagnostics by making it possible to visualize tissue at remote internal sites in a minimally invasive fashion. The most common form employs an imaging fiber bundle to provide high quality white light reflection imaging. Laser scanning confocal reflection and fluorescence endoscopes also exist and can provide 3D cellular resolution in tissues. Confocal endoscopes are now becoming available commercially (Optiscan Ltd, Australia, Lucid Inc, Rochester) and are being applied in a number of clinical trials for cancer diagnosis. Multiphoton excitation based endoscopes has attracted significant attention recently. There were a number of advances, including fiber delivery of excitation pulses, miniature scanners, double clad fibers for efficient signal collections, etc. Thus, just like MPM has proven to be a powerful tool in biological imaging, multiphoton endoscopes have great potentials to improve the capability of the existing laser-scanning optical endoscopes. It is quite obvious that a compact, fully electronically controlled, femtosecond system seamlessly integrated with fiber optic delivery is essential for multiphoton endoscopy in medical diagnostics, particularly to biomedical experts who are not trained in lasers and optics.
Perhaps the most promising and successful area in biomedical imaging that showcases the unique advantage of multiphoton excitation is imaging deep into scattering tissues. In the past 5 to 10 years, MPM has greatly improved the penetration depth of optical imaging and proven to be well suited for a variety of imaging applications deep within intact or semi-intact tissues, such as demonstrated in the studies of neuronal activity and anatomy, developing embryos, and tissue morphology and pathology. When compared to one-photon confocal microscopy, a factor of 2 to 3 improvement in penetration depth is obtained in MPM. Nonetheless, despite the heroic effort of employing energetic pulses (˜μJ/pulse) produced by a regenerative amplifier, MPM has so far been restricted to less than 1 mm in penetration depth. One promising direction for imaging deep into scattering tissue is to use longer excitation wavelength. Although there is little data for tissue scattering beyond 1.1 μm, the available data at shorter wavelengths clearly indicates the general trend that tissue scattering reduces as one uses longer excitation wavelength. Longer excitation wavelengths are an obvious choice for imaging deep into scattering tissues.
A concern for longer wavelength excitation is water absorption, which is typically the dominant contribution to tissue absorption at near IR. However, a careful examination showed that water absorption at 1270 nm is only approximately twice as high as that between 960 and 1000 nm, a spectral region where multiphoton excitation and imaging has been routinely performed in the past. Although the “diagnostic and therapeutic window,” which is in between the absorption regions of the intrinsic molecules and water, extends all the way to ˜1300 nm, previous investigations involving multiphoton imaging are almost exclusively carried out within the near IR spectral window of ˜700 to 1100 nm, constrained mostly by the availability of the excitation source. Currently, there are only two femtosecond sources at the spectral window of 1200 to 1300 nm, the Cr:Forsterite laser and the optical parametric oscillator (OPO) pumped by a femtosecond Ti:Sapphire (Ti:S) laser. In terms of robustness and easy operation, both sources rank significantly below the Ti:S laser. Thus, the development of a reliable fiber source tunable from 1030 to 1280 nm will open up new opportunities for biomedical imaging, particularly for applications requiring deep tissue penetration.
Shortly after the inception of MPM, mode-locked solid state femtosecond lasers, most commonly the Ti:S lasers, have emerged as the favorite excitation sources to dominate the MPM field today. When compared to earlier ultrafast lasers, e.g., ultrafast dye lasers, the Ti:S lasers are highly robust and flexible. The concurrent development of the mode-locked Ti:S lasers was perhaps the biggest gift for MPM and enabled MPM to rapidly become a valuable instrument for biological research. Nonetheless, the cost, complexity, and the limited potential for integration of the bulk solid state lasers have hampered the widespread applications of MPM in biological research. The fact that a disproportionate number of MPM systems are located in physics and engineering departments, instead of the more biologically oriented institutions, reflects at least in part the practical limitations of the femtosecond pulsed source. Obviously, the requirement of a robust, fiber delivered, and cheap source is even more urgent for multiphoton endoscopy in a clinical environment.
For these reasons, previous work using SSFS below 1300 nm was performed at soliton energies either too low or too high (by at least an order of magnitude) for many practical applications, such as multiphoton imaging, where bulk solid state lasers are currently the mainstay for the excitation source.
The present invention is directed to overcoming these and other deficiencies in the state of the art.